1. Field of the Invention
This invention relates generally to magnetic resonance imaging (MRI) utilizing nuclear magnetic resonance (NMR) phenomena. It more particularly relates to the drive circuits (and control thereof) main or polarizing magnetic field B.sub.o of the typical for gradient coils and/other similar electromagnet coils utilized to supplement the nominally static MRI main magnet structure.
This invention may be considered as generally related to the following prior, commonly assigned, U.S. patent application:
Ser. No. 07/786,828, filed Nov. 1, 1991 (now U.S. Pat. No. 5,227,723) naming Messrs. Kaufman, Carlson and Gran as inventors and entitled "Gradient Driver Control In Magnetic Resonance Imaging."
2. Related Art
Magnetic resonance imaging systems are now commercially available from a number of sources. There are, in general, several techniques known to those in the art. Some exemplary MRI techniques are set forth, for example, in commonly assigned issued U.S. Pat. Nos. 4,297,637; 4,318,043; 4,471,305 and 4,599,565. The content of these issued patents is hereby incorporated by reference.
In all MRI systems, a main polarizing magnet structure is used to provide a substantially uniform homogeneous magnetic field within a patient image volume along a predetermined axis (e.g., the z axis of the usual x,y,z orthogonal coordinate system). When precisely controlled gradient magnetic fields are superimposed within the image volume along various different axes, the detectable NMR characteristics of NMR nuclei can be spatially encoded (in conjunction with suitable RF nutation pulses) so as to produce RF responses that can be processed to provide two dimensional arrays of display pixel values representing two and/or three dimensional representations of NMR nuclei within the patient volume. However, the accuracy of the MRI process is directly related to the degree of homogeneity in the static field and the degree of linearity in imposed gradient fields along selected axes (the gradient field being a constant as a function of position with respect to other orthogonal axes within the patient volume). To the degree that these desired goals of static field homogeneity and of gradient field linearity along precisely selected axes fail to be attained, then inaccuracies and/or artifacts in the resulting displayed image can be expected. Accordingly, considerable effort previously has been expended toward the ends of either attaining these goals or toward correcting or compensating for deviations from such goals.
Some MRI systems have main polarizing magnet structures that include permanent magnets and/or iron elements in the relevant main magnet magnetic circuit. For example, a relatively low field open architecture main magnet is employed in the MRI system design described by related U.S. Pat. No. 4,829,252. In such structures, the iron components exhibit remanent magnetization and hysteresis effects as a result of their past magnetization history. This can cause a number of undesirable effects such as image distortion, intensity variations, artifacts of various sorts, etc.
Remanent magnetization and hysteresis effects are especially troublesome with respect to the rapid sequential changing energization of gradient coils. For example, the usual phase encoding gradient coils (e.g., oriented to produce a gradient in the y axis dimension) typically are sequenced in 256 steps from a drive pulse of maximum positive value to a maximum negative value (e.g., over 256 successive MRI data gathering cycles). The first large pulse will leave a relatively large remanent magnetization throughout the next 127 successively smaller magnitude positive drive pulses. However, as the drive pulse polarity reverses the remanent gradient will also be caused to be reversed and will then stepwise increase in magnitude throughout the remainder of the complete imaging sequence to a maximum negative value. Furthermore, even for gradient axes which do not change in magnitude or polarity during the entire imaging sequence (e.g., as along the x axis which is typically energized during NMR RF signal readout), a remanent gradient of some sort will remain to possibly distort a subsequent imaging sequence.
Ambient or other environmental changes can also cause undesirable changes in the magnetic field for an MRI system. For example, local changes in the earth's magnetic field or local magnetic field changes induced by local movements of large magnetically permeable masses (e.g., elevators, locomotives, etc.), by magnetic fields caused by the passage of large local electrical currents (e.g., as in the drive circuits of elevators, trolley cars, trains, subways, etc.), by ambient temperature changes and related changes in magnetic circuit properties, by hysteresis effects in magnetically coupled bodies, and the like, are all potential sources of undesired deviations from the optimum spatial distribution of magnetic field orientation and strength within the patient image volume of an MRI system. Such deviations may change throughout any given imaging procedure or over the historical period of system installation at a given site (thus perhaps impairing the ability to accurately compare images taken at the same site at widely separated intervals of time) or ambient conditions as compared between different sites. Currently, substantial efforts are required during installation of an MRI system at a particular site in an attempt to minimize such difficulties. Extra care must currently also be taken to assure quality and repeatability in magnet production for MRI systems. Many special processing steps or other precautions are often required to provide reliability and image quality of sufficiently high standards in view of these ongoing problems. Accordingly, there is considerable need for a more comprehensive and efficient technique to minimize the possible adverse effects of such potential problems.
Another potential source of similar problems are eddy currents generated by rapidly changing magnetic gradient fields in surrounding electrically conductive materials. Associated with each attempted change in the magnetic gradient flux will be the generation of eddy currents which, in conformance with Lenz's law, will produce a magnetic field which opposes the attempted change in the gradient field. Accordingly, it has long been known that some kind of eddy current compensation must be included in the drive current supplied to a gradient coil.
As earlier noted, consistent, reliable operation of a magnetic resonance imaging system relies strongly on the creation of ideal gradient flux pulses inside the volume to be imaged. Nearby electrically conductive structures inherently support eddy current loops when exposed to the rapidly switched gradient fields and these result in distortions to the desired spatial distribution of magnetic flux. Such eddy currents, located in nearby metallic structures, decay in a manner that is characteristic of a collection of somewhat different exponential time constants. If not compensated, the time variation thereby produced in the net magnetic flux actually present at the patient image volume would be sufficiently severe to distort section profile and end-plane resolution of the imaging system.
As a consequence, magnetic resonance images have long used some kind of compensation to reduce the effects of such secondary "eddy" currents. The most common prior technique is an open loop feedback system whereby the gradient flux demand pulse is purposely overdriven (e.g., "pre-emphasized" in a predetermined and pre-calibrated waveform). Determining the exact characteristics of such overdriving for a particular installation site presently requires a considerable and lengthy effort. Hopefully, once this laborious process has been completed, the open loop control system will overdrive the gradient coil in just the right manner to thereafter anticipate the induced eddy currents and to result in a net actual flux field that approximates the ideal. However, not only does this kind of system setup consume considerable time initially and thereafter (in a maintenance mode), it is virtually impossible to find one predetermined overdrive specification that will properly compensate for eddy currents under changing operational conditions. For example, if the magnet structure is a cryogenic super conducting magnet, then as the cryogen boils off, the temperature of various metallic conductor elements may vary which, in turn, causes a change in resistivity and a change in the time response of the eddy current subsystems. Furthermore, spatial variations in eddy current fields often do not exactly track the gradient coil flux field (e.g., spatial dependence may also vary as a function of temperature and other changes in the system). In short, it is virtually impossible for a simple open loop compensation system to exactly correct for eddy current effects. A typical overdrive compensation involves a current overshoot of approximately 20% with a decay to an asymptotic value involving two or three time constants plus a similar undershoot when the drive pulses turn "off" with a similar multi-time constant decay to the asymptotic zero current state.
A less common but somewhat better technique for reducing adverse eddy current effects is to wind a shield coil around the gradient coil. Although this may substantially eliminate the effect of some eddy current fields (e.g., those induced in the aluminum cryogenic container), it occupies a considerable additional portion of the available magnet bore space thus substantially decreasing access to the image volume while adding substantial cost, weight, etc., to the overall MRI system.
As a part of the lengthy setup procedure now required for installation of an MRI system at a particular site, considerable effort is often given to centering the gradient coils in an attempt to avoid asymmetric eddy current effects. If the eddy currents are substantially asymmetric, then there may be no technique known in the prior art for adequately compensating them.
Prior gradient coil drive circuitry typically utilizes a switched power supply with an analog feedback loop used to provide pulse width modulation of the power supply switching. However, such analog feedback for current control in switching power amplifiers may facilitate undesirable drift and/or noise in gradient power amplifiers and may provide a relatively restricted dynamic range. To the extent that such analog circuits obtain analog feedback by a current sensing resistor connected in series with the gradient coil, significant power losses are also incurred in conventional gradient drive controllers. Currently, even those having "digital inputs" are believed to have digital to analog converters so as to produce the usual analog control signal for use in the usual analog feedback loop.
The earlier referenced related application Ser. No. 07/786,828 describes several improvements for gradient drive control circuits in magnetic resonance imaging systems which substantially alleviate or at least improve many of the above-mentioned problems. Perhaps the greatest such improvement is the provision of a closed-loop real-time feedback control for the gradient coil drivers in an NMR system. Here, the actual net gradient flux (e.g., including hysteresis and eddy current effects) is monitored during the imaging procedure on a real-time basis. These inputs are then presented to an integrating feedback controller for each gradient coil control channel so as to modulate the gradient coil current as needed, on a real-time basis, to maintain whatever desired gradient value is then indicated from the main MRI system control (e.g., typically a fixed gradient flux magnitude and polarity for a given length of time).
The flux monitoring coils are preferably patterned after the gradient coils (e.g., properly oriented sets of "saddle" coils of the Golay variety). Helmholtz and/or Maxwell coil configurations (e.g., see U.S. Pat. No. 4,755,755--Carlson) of pancake D-shaped coils and the like may be used if the typical open-architecture permanent magnet structures are employed (e.g., see U.S. Pat. No. 4,829,252--Kaufman). However, the gradient flux sensing coils must, of course, be located and/or dimensioned so as to occupy a different physical volume than the gradient flux generating coils. For a solenoidal cryogenic superconducting magnet structure, the gradient flux sensing coils may preferably have a slightly reduced radius as compared to the gradient flux generating coils so that they may be closely located just inside the usual gradient coil structure. It should be noted that in presently preferred exemplary embodiments, typically only single turn coils are utilized for the flux sensing coils. Since the conductor of the flux sensing coil does not need to be extensive (e.g., 0.010 to 0.020 inches thick by 0.025 to 0.5 inches wide copper strip), only a slight further incursion is necessitated into the magnet bore space so that no substantial further restriction on patient image volume or access to same is necessitated. To obtain proper balance between the mutual inductance (e.g., magnetic coupling) between the flux sensing coils and the flux generating coils on the one hand and the eddy current/hysteresis flux on the other hand, it may be necessary to slightly offset the flux sensing coils with respect to the flux generating coils (thereby somewhat degrading the mutual inductance between these two sets of coils).
Although the preferred exemplary embodiment utilizes flux sensing coils which are a substantial replica of the gradient flux generating coils, the net magnetic flux within the patient imaging volume may be monitored by other arrangements. For example, Hall effect probes, flux gate magnetometers, conventional pickup loops, SQUIDS, electron magnetic resonance detectors (like the ferrites in YIG oscillators) etc., may be strategically located with outputs appropriately combined so as to detect changes from an initial or desired magnetic flux state.